Biochips or DNA chips include a flat carrier, on one side of which at least one spot array, that is to say a gridlike arrangement of analysis positions, is present. The spots contain probe or catcher molecules, for example oligonucleotides, immobilized on the carrier surface. Target molecules, for example DNA fragments, contained in an analyte solution applied to a spot couple to the catcher molecules. The conversion of such coupling or binding events into detectable signals is effected by means of optical, piezoelectric, electrochemical, calorimetric or impedance-spectroscopic methods.
In the case of an impedance-spectroscopically readable DNA chip disclosed in DE 196 10 115 C2, an interdigital electrode arrangement is present on a sensor area, catcher molecules being immobilized on the electrodes and the areas arranged between the electrodes. The coupling of target molecules to the catcher molecules leads, e.g. on account of charge changes, to a change in the alternating electric field generated by the electrodes or generally to a change in an electrical property in the vicinity of the electrodes, e.g. the impedance. A measurement of an impedance change can be carried out by way of a, for example, two-pole interdigital electrode arrangement in which the electrodes are formed from a plurality of partial electrodes.
What is problematic with the last manner of detecting binding events is that the dimensions of the electrode structures differ by orders of magnitude from molecular dimensions. With a technical outlay that is still tenable, it is possible to produce electrodes whose width and spacing, taken together, have a value L (=width+spacing) of approximately 2 to 20 μm and a height of approximately 0.1 to 0.5 μm.
The impedance-spectroscopically detectable range of the electric field of such an electrode arrangement extends approximately 1 to 5 L (=2 to 100 μm) beyond the carrier surface or the planar plane spanned by the electrode arrangement. By contrast, a catcher molecule having 100 base pairs, for example, has a length of only approximately 30 nm. The influence of binding events in a monomolecular layer of catcher molecules that is immobilized on the sensor surface or the electrodes on the electric field is correspondingly low, particularly when only few binding processes take place. The publication “Nanoscaled interdigitated electrode arrays for biochemical sensors”, P. van Gerwen et al, Sensors and Actuators B 49, 1998, 72-80, proposes, for solving the problem discussed, approximating the dimensions of electrode structures to the dimensions of DNA target molecules, electrode structures with partial electrodes being sought whose widths and mutual spacings lie approximately in the range of 250 to 500 nm. However, such dimensions are associated with an increased production outlay.
Furthermore, WO 98/19153 A1 discloses a sensor for biochemical applications which contains electrodes embedded in a conductive polymer. In this case, the conductive polymer is in contact with the analyte in which a biochemical process takes place as a result of alternating-current influencing. In this case, process changes by way of the conductive polymer are forwarded as impedance changes to the electrode system and detected.
The sensitivity of a sensor chip constructed in this way is problematic. Moreover, embedding or coating the electrodes in a conductive polymer is complicated, so that the biosensor described is not practically suitable.